Apparatus and method for medical imaging

ABSTRACT

The present invention discloses an apparatus for use in medical imaging including a readout circuit having an input for receiving a detection signal corresponding to a photon hitting a radiation detector, wherein the readout circuit is adapted to output, in response to receiving said detection signal, a pulse signal having a leading edge encoding a time-stamp of said photon and a width encoding the energy of said photon. A method of reading out detection signals from a radiation detector of a medical imaging apparatus is also provided.

BACKGROUND OF THE INVENTION

The present invention relates to an apparatus for use in medical imagingcomprising a readout circuit having an input for receiving a signalcorresponding to a photon hitting a radiation detector. The applicationfurther relates to a detector assembly and to a combined PET-CTapparatus as well as to a method according to the preamble of claim 27.

Medical imaging techniques employing radiation detectors for detectingphotons for example comprise Positron Emission Tomography (PET), X-raycomputed tomography (CT), Single Photon Emission Computed Tomography(SPECT) and γ cameras. While the invention is not limited to anyspecific type of medical imaging, for illustrative purposes we shallexplain the invention with specific reference to PET.

PET is a nuclear medicine tomographic imaging technique using γ rays. Toconduct a scan, a short-lived radioactive tracer isotope which decays byemitting a positron and which has been chemically incorporated into ametabolically active molecule is injected into the patient, typicallyinto the blood circulation. There is a waiting period while themetabolically active molecule becomes concentrated in the tissues ofinterest. Thereafter, the patient is placed in an imaging scanner.

As the radioisotope undergoes positive beta decay, it emits a positronwhich after a short travel encounters and annihilates with an electron,thereby producing a pair of γ-photons having an energy of 511 keV andusually traveling in opposite directions. The γ-photons are detected bya radiation detector typically comprised of a scintillator and anassociated photodetector. The signal from the photodetector must then bereadout by a suitable readout circuit. PET depends on simultaneous orcoincident detection of the pair of γ-photons. Photons which do notarrive in pairs (i.e. within a few nanoseconds) are ignored.

Since most of the γ-photons are emitted at 180 degrees from one anotherduring electron-positron annihilation, the source of radiation can belocated along a straight line connecting the two radiation detectorsites at which coinciding hits are detected. If the response of theradiation detector and the corresponding reader circuit is fast enough,it is moreover possible to calculate the location of the radiationsource on said line from a difference in the arrival times. This methodis called “time-of-flight” (TOF) measurement. However, this wouldrequire a time resolution of the measurement in the picosecond rangewhich is currently difficult to achieve. Instead, typically statisticsare collected from tens-of-thousands of coincidence events and equationsfor the total activity of each parcel of the tissue of interest alongthe above mentioned straight line is solved, such that a map ofradioactivities can be constructed and outputted. Clearly, iftime-of-flight information is available, the statistics needed forproducing a high quality image would be smaller, and accordingly, asmaller dose of radioactive tracer isotope could be used which would behealthier for the patient. A prime goal of the above mentioned apparatusis therefore to provide a very fast readout channel allowing formeasurements in the picosecond rather than nanosecond range.

FIG. 1 schematically shows an electronic readout channel according toprior art which is currently used in PET-systems. The readout channel ofFIG. 1 comprises and avalanche photo diode (APD) array element 10receiving light from a scintillator (not shown) upon a photon hittingthe scintillator. The APD array element outputs an analog signal whichis processed by a preamplifier 12 and a shaper 14 and is then inputtedinto an analog-to-digital (A/D) converter 16. The A/D converter 16 isclocked at high frequency (typically 40 to 100 MHz) and outputs adigital signal. The digital signal is then transferred to a FIFO buffer18 and to a coincidence processor 20 for detecting coincidence of twoγ-photons which can be assigned to a common electron-positronannihilation. The digital samples are processed by an algorithmimplemented in a digital filter that determines the time-stamp and asignal of the amplitude. The architecture of FIG. 1 continuously samplesanalog signals and produces large data sets that due to their large sizeare generally processed off-line further down the image processingchain.

The prior art readout architecture of FIG. 1, however, has the followingdisadvantages.

-   1. Retrieving time information for sampled data requires additional    hardware and algorithms that increase processing time and hardware    costs.-   2. The precision of the sampling of the signals depends on the    signal shape. Shape variations lead to a decrease of precision in    the evaluation of the timing of the corresponding event, in    particular since PET events are randomly distributed over time and    thus uncorrelated with the sampling clock cycle.-   3. To obtain a higher timing precision and event counting rate, the    clock frequency of the sampling A/D converter must increase which    might lead to significant data overflow problems.-   4. Modern PET scanners and in particular whole-body-PET scanners    with a large field of view use increasingly more detector rings    containing highly segmented arrays of scintillating crystals each    being as small as 2×2 mm², such that the number of required    front-end electronic readout channels easily increases to more than    one hundred thousand. For such a large number of channels the cost,    power dissipation and complexity of signal processing using the    readout architecture of FIG. 1 becomes extremely demanding.

As mentioned above, the invention is not limited to PET-imaging, andaccordingly, in FIG. 2 a prior art readout architecture for X-raydetection is shown. In the shown front-end electronics of an X-rayCT-scanner system an APD-array element 10 collects light from ascintillator (not shown) hit by an X-ray photon and generates an analogcurrent signal. The analog current signal is amplified by a currentamplifier 22, and the amplified current is integrated by an integrator24. The integral of the current received in a given time frame isproportional to the number of X-ray photons hitting the detector in saidtime frame. The integrated signal is then digitized by an A/D converter26 and outputted to an image reconstruction processing logic. Note thatin the context of this application, the term “logic” refers to any kindof hardware, software or combination thereof providing the respectivelogic function.

PET-scans and CT-scans are preferably performed simultaneously sincethey give in combination both anatomic and metabolic information. Thatis, modern PET-scanners are available with integrated high-endCT-scanners. Since the two types of scans can be performed sequentiallywithout the patient having to move between the scans, the two sets ofimages are precisely registered such that areas of the PET imaging canbe precisely correlated with the anatomy provided by the CT-images. Notehowever, that the readout technologies of FIGS. 1 and 2 are incompatiblewith one another such that it is not possible to perform a simultaneousPET-CT co-registration with a common detector head.

In other prior art PET systems the photon arrival time is detected usinga constant fraction discriminator (CFD) that extracts the precisetime-stamp of photon arrival for coincidence and time-of-flightmeasurement. CFDs are for example known for the detection ofscintillator pulses having identical rise times which are much longerthan the desired temporal resolution. Accordingly, it is not sufficientto use a simple threshold triggering which would introduce a dependenceof signal peak height and trigger time, an effect which is called “timewalk” and is described more in detail below. Instead, according to CFDtriggering is not performed on a fixed threshold but on a constantfraction of the total peak height yielding trigger times independentfrom peak heights.

An example of a CFD circuit is shown in FIG. 3. A CFD circuit used forPET-applications is for example described in detail in “An Analog SignalProcessing ASIC for a Small Animal LSO-APD PET Tomograph”, V. Ch.Spanoudaki, D. P. McElroy and S. I. Ziegler, Nuclear Instruments andMethod in Physics Research, Section A, in press, such that a detailedexplanation is omitted.

However, a readout circuit based on a CFD has two main disadvantages.Firstly, the integration of a fast CFD circuit in a monolithic CMOS thatis now currently done for electronics system turns out to be difficultfor timing precision of better than 100 ps. Also, the CFD circuits arequite expensive if produced in the necessary amount. Secondly, aCFD-based readout architecture only provides a time-stamp but not thesignal amplitude, such that PET events cannot be reliably distinguishedfrom background events and Compton events cannot be reconstructed.

Accordingly, it is an object of the current invention to provide amethod and an apparatus for use in medical imaging that overcome theabove mentioned drawbacks.

SUMMARY OF THE INVENTION

This object is achieved by an apparatus according to claim 1 and amethod according to claim 27. Preferable embodiments are defined in thedependent claims.

According to the invention, the readout circuit of the apparatus isadapted to output in response to receiving the detection signal a pulsesignal having a leading edge encoding a time-stamp of the photon and awidth encoding the energy of the photon. That is, the information ofinterest, namely the time-stamp and the energy is encoded in a simpledigital pulse signal. Accordingly, the invention compares favorably withan apparatus based on digitizing detection signals where large amountsof data are produced, and it is at the same time is superior to CFDtechniques in providing both time-stamp and energy information.Moreover, such a pulse signal can be easily and rapidly processed by atime-to-digital converter (TDC), as will be explained below, such thatimages can be generated in real time and that events can be recorded atextremely high rates.

In a preferred embodiment, the apparatus further comprises means forcorrecting the time-stamp of the pulse signal based on the width of thepulse signal. Said means may be configured to estimate a time-walk ofthe pulse signal from the width of the pulse signal and to subtract thetime-walk from the time-stamp, as will be explained in greater detailbelow.

In a preferred embodiment, the readout circuit is adapted for encodingthe time-stamp and energy of one or more types of photons selected amonga group consisting of: γ-photons as generated during electron-positronannihilation, X-ray photons of a wavelength suitable for X-ray CT,photons in the visible range, and γ-photons as emitted by radiopharmaceuticals suitable for γ camera imaging or SPECT. That is, theapparatus cannot only be adapted for a particular type of photonsuitable for medical imaging, but it can moreover be simultaneouslyadapted for two different types of photons. For example, the apparatusmay be suitable for reading out both detection signals corresponding toCT X-ray photons and detection signals corresponding to PET photons.Accordingly, in such a configuration the apparatus allows for a singledetection head or front-end which is capable of detecting both, CT andPET signals.

In a preferred embodiment, the readout circuit is a discriminatorcircuit configured to compare the detection signal to a threshold value,where the leading edge of the pulse signal corresponds to the time thedetection signal first exceeds said threshold value and the trailingedge of the pulse-signal corresponds to the time the detection signaldrops below said threshold value. As will be described in more detailbelow, such a discriminator circuit can be manufactured at low cost andis capable of operating at very high speed sufficient for time recordingwell in the picosecond range.

In a preferred embodiment, the apparatus comprises a filter circuit forfiltering the pulse signals according to their width. In particular, thefilter may comprise at least one gate-delay circuit comprising anAND-gate to which a fraction of the pulse signal and a delayed fractionof said pulse signal is fed, which is delayed by a predeterminedtime-delay. Such a filter is simple to implement and at the same timeeffective in discriminating between signals corresponding to differenttypes of events or different types of photons.

In a preferred embodiment the filter circuit comprises a PET-filterconfigured to pass pulse signals having a width corresponding to theenergy of γ-photons emitted during electron-positron annihilation. Inone embodiment, the PET-filter blocks pulse signals having a widthcorresponding to an energy of less than 350 keV, and preferably, lessthan 400 keV.

Alternatively or additionally, the filter circuit may comprise aCT-filter, configured to allow pulse signals having a widthcorresponding to the energy of X-ray photons suitable for CT to becountered by a counter. With these types of filters, it is possible touse the same readout circuit for both CT and PET events and todistinguish between the signals downstream. In particular, this allowsfor a combined PET-CT apparatus with a common detector head for both, CTand PET signals.

In a preferred embodiment, the apparatus comprises a filter fordistinguishing pulse signals corresponding to two or more detectionsignals corresponding to individual X-ray photons that are overlappingin time. In particular, the filter may comprise a number of n gate-delaycircuits as defined above connected in parallel, where n is an integergreater or equal than 2, wherein the predetermined time-delays of thegate-delay circuits respectively correspond to an expected width of thepulse signal for one, two, . . . , n overlapping X-ray detectionsignals. This filter will allow for a so-called pile-up correctionoccurring when CT photons at a given detector segment overlap in time.Such pile-up will be likely to occur with an increasing X-ray intensity.This filter allows to avoid the corruption of the X-ray image due tosignificant overlapping of X-ray photons at higher X-ray intensities.

Preferably, the apparatus further comprises a time-to-digital (TDC)converter connected to receive the pulse signal and configured tomeasure the arrival time of both, the leading and trailing edge of thepulse signal. The TDC may further be configured to store the leadingedge time measurement and the pulse width in a local channel FIFOregister.

The invention further relates to a detection assembly for a medicalimaging apparatus, and in particular, a PET-apparatus or a combinedPET-CT-apparatus comprising a multitude of apparatuses according to oneof the embodiments described above. The detection assembly may comprisea multitude of detector elements, each detector element comprising anarray of radiation detectors providing a number of channels foroutputting detection signals and a corresponding number of apparatusesas defined above providing electronic channels for reading out thedetection signals.

Preferably, the detection assembly comprises a control unit configuredor programmed for detecting coincidence of detection signals.Preferably, the control unit is further configured or programmed tocalculate the location of an electron-positron annihilation using adifference in time-stamp of the coinciding events.

In a preferred embodiment of the detection assembly, some or all of themultitude of detector elements are connected to a common referenceclock.

Finally, one aspect of the invention relates to combined PET-CTapparatus comprising an X-ray generator suitable for CT and a detectionassembly according to one of the embodiments described above.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a prior art readout architecture for a PET apparatus basedon sampling amplified analog photodetector signals with an A/C converterand processing the sampled signals with a digital processor whichextracts time-stamp and energy amplitude.

FIG. 2 shows a prior art readout architecture for measuring an X-rayphoton intensity based on a current signal integration of thephotodetector sensing light produced by a scintillator placed in frontof the photodetector during one CT image frame.

FIG. 3 shows a prior art readout architecture for a PET-apparatus basedon a CFD discriminator and a pulse peak detector.

FIG. 4 shows a schematic illustration of a combined PET-CT apparatusaccording to an embodiment of the invention.

FIG. 4A is a block diagram showing a combined PET-CT readoutarchitecture employing a pulse modulation discriminator and a TDCaccording to an embodiment of the invention.

FIG. 5 shows an example of a preamplifier.

FIG. 5A shows a diagram of ENC as a function of the input capacitancefor two different input transistor bias currents calculated for an 800nA feedback bias current.

FIG. 6 shows an example of the fast pulse-modulation discriminatorintegrated in a monolithic CMOS technology.

FIG. 7 shows an example of the input stage of the discriminator of FIG.6.

FIG. 8 shows a schematic timing diagram related to the pulse widthmodulation discriminator of FIG. 6 when operating in PET mode.

FIG. 9 shows an output waveform of experimental measurements obtained atthe output of the pulse width modulation discriminator of FIG. 6.

FIG. 10 shows a relation between time-walk and pulse width as obtainedwith the discriminator circuit of FIG. 6.

FIG. 11 shows a schematic timing diagram for the discriminator circuitof FIG. 6 operating in CT mode.

FIG. 12 is a circuit diagram for a filter for discriminating X-raypulses overlapping in time.

FIG. 13 is a circuit diagram showing filters for distinguishingCT-signals from PET-signals based on signal pulse width.

FIG. 14 is a conceptional drawing showing overlapping pulse signals.

FIG. 15 shows a TDC-circuit architecture based on common clockreference.

FIG. 16 is a block diagram illustrating a PET readout architectureemploying TDCs with local time measurements based on global timereference.

FIGS. 17 and 18 shows histogram plots of emission spectra obtained for⁵⁷Co, ²²Na, respectively obtained with the readout architecture of thepresent invention.

DESCRIPTION OF THE PREFERRED EMBODIMENT

For the purposes of promoting an understanding of the principles of theinvention, reference will now be made to the preferred embodimentillustrated in the drawings, and specific language will be used todescribe the same. It will nevertheless be understood that no limitationof the scope of the invention is thereby intended, such alterations andfurther modifications in the illustrated devices and method and suchfurther applications of the principles of the invention as illustratedtherein being contemplated as would normally occur now or in the futureto one skilled in the art to which the invention relates.

In FIG. 4 a combined PET-CT-apparatus 28 according to an embodiment ofthe invention is schematically shown. The PET-CT-apparatus 28 comprisesa plurality of detector rings 30, one of which is shown in FIG. 4. Thedetector ring is surrounding a body of a patient 32 which is subject toa combined PET- and CT-examination. The detector ring 30 is comprised ofa multitude of detector elements 34, of which only two are shownschematically in FIG. 4.

Each detector element comprises a scintillator array 36 and aphotodetector array 38 located adjacent to the scintillator array 36. Inthe shown embodiment, the photodetector array 38 is comprised of anarray of avalanche photodiodes.

Each of the photodetector arrays 38 is connected with a front end orreadout circuitry 40 providing a readout architecture to be described ingreater detail below. A time-to-digital converter (TDC) 42 is connectedto each front-end 40. All of the TDCs are connected to a coincidenceprocessor 44 which in turn is coupled with a PET-CT reconstructionprocessor 46. The combined PET-CT apparatus 28 of FIG. 4 furthercomprises a rotating X-ray generator 48 for generating X-rays suitablefor CT.

In the following, the general function of the combined PET-CT apparatus28 of the embodiment of FIG. 4 is described. Before the patient 32 isplaced in the detector ring 30, a short-lived radioactive tracer isotopethat has been chemically incorporated into a metabolically activemolecule is injected to the patient 32. There is a waiting period whilethe metabolically active molecule, such as fluorodeoxyglucose, becomesconcentrated in the tissues of interest. Upon the decay of theradioactive tracer isotope a positron is emitted. The positron willafter a travel of at most a few millimeters encounter an electron andannihilate with the same at an annihilation point indicated at 48 inFIG. 4.

Upon the electron-positron annihilation, two 511 keV γ-photons areemitted under an angle of 180° from one another. The γ-photons will bedetected at two different detector elements 34. Using the readoutcircuitry of the front-end 40 and the TDC 42, the arrival time, alsocalled time-stamp, and the energy of the detected photons 50 aredetermined. The energy of the measured photon can be used to determinewhether the photon has the expected energy and is therefore due to theelectron-positron annihilation and not for example due to backgroundnoise. The time-stamps of the detection events are used by thecoincidence processor 44 to determine whether two detected events arecoinciding in time such that they can be assigned to a commonannihilation process.

Note that the two photons 50 would have exactly the same time-stamp onlyif the annihilation point was exactly at the center of the detector ring30. With reference to FIG. 4, the annihilation point 48, however, iscloser to the left detector element 34 than to the right detectorelement 34 such that the time-stamp of the photon detected at the leftdetector 34 will be slightly earlier than the time-stamp of the photondetected by the right detector element 34. Such events will still bedetected as coinciding by the coincidence processor 44. In fact, thedifference in time-stamp is used by the PET-CT reconstruction processor46 to calculate at which location along the line of propagation of thephotons 50 the annihilation took place. Such a type of measurement iscalled “time-of-flight” (TOF) measurement for obvious reasons. From thecoincidence information and time-of-flight information the PET-CTreconstruction processor 46 can generate a 3 D image of positronemission events which can for example be used in clinical oncology forthe medical imaging of tumors and the search for metastases.

Simultaneously with the PET scan a CT scan is performed using therotating X-ray generator 48. The X-ray photons will be detected by thesame detector elements 34 as the PET-γ-photons, such that the samescintillator 36, photodetector array 38 and most of the front-end 40 canbe used for the X-ray imaging simultaneously as well. Since the readoutarchitecture provides the energy of the event in addition to itstime-stamp, and since the X-ray photons are on the order of 100 keV andtherefore roughly a fifth of the energy of the PET-γ-photons, thereadout architecture can discriminate between the two types of events,as will be shown in greater detail below. The PET-CT reconstructionprocessor can then use the information from the PET and the CT scans ofthe same patient to thereby generate two sets of images which areprecisely registered so that areas of abnormality on PET-imaging can beperfectly correlated with the patient's anatomy imaged by CT.

As is clear from the above discussion, the electronic readout channelsused for the PET-CT apparatus 28 must have a number of specificproperties. In particular, the readout architecture must be capable ofoperating fast enough such as to produce data of a precision in thepicosecond range to allow for TOF measurements and also fast enough tomanage high photon rates. Also, since there are a multitude of detectionchannels (a number of 100,000 can easily be reached), the readoutarchitecture should be low priced and have limited power dissipation. Inaddition, the amount of data produced during readout should be kept lowsuch that it is possible to generate the images “on the fly”. Due todata processing capacity, if the amount of data produced is too large,the only possibility is to store the data and construct the imagesoffline.

In order to meet the above prerequisites, an aspect of the inventiondeals with a novel readout architecture which will be described in thefollowing in more detail.

In FIG. 4A, a block diagram of a detector element 34 and a correspondingreadout architecture is shown. The detector element 34 provides for 32channels, of which in FIG. 4A only four are shown. The detector element34 comprises a 32 channel LSO-crystal array attached to a 32 channelAPD-array 38. The output of each APD-element of the APD-array 38 isconnected to an electronic readout channel. Each electronic readoutchannel comprises a preamplifier 52, a pulse modulation discriminatorcircuit 54 and a PET-CT-filter circuit 56 located downstream of thepulse modulation discriminator 54. Each pulse modulation discriminator54 receives an analog signal corresponding to a preamplified detectionsignal from a photon hitting the scintillator 36 and generates therefrom a square pulse signal having a leading edge encoding a time-stampfor that photon and a width encoding the energy of the photon. Since thewidth of the pulse signal is modulated according to the energy of thedetected photon, the pulse signal may be termed a “pulse widthmodulated” signal.

The filter 56 discriminates the pulse signals outputted from the pulsemodulation discriminator 54 by their pulse widths, that is, by theenergy of the corresponding photon. In the embodiment shown in FIGS. 4and 5, the filter 56 discriminates between pulses arising from X-rayphotons and pulses arising from PET-γ-photons. Pulses corresponding toX-ray photons are channeled to a counter 58 which is suitable forcounting the number of X-ray photons detected in the respective channelwithin a given time-frame, thus providing one pixel of an X-ray image.The filter further directs pulse signals corresponding to PET-γ-photonsto a TDC-circuit 42 in which the arrival time of both, the leading andtrailing edge of the pulse signal are measured. In a preferredembodiment, each TDC 42 is receiving a global reference clock signal 60.This global clock reference signal will allow to compare time-stampsfrom different channels, and in particular, from channels of differentregions of the detector ring 30, which is important for coincidence andTOF-measurements.

In FIG. 5 a circuit diagram of one channel of the 32 channelpreamplifier 52 of FIG. 4A is shown in detail. In the shown embodiment,the preamplifier circuit 52 is implemented in a standard 0.25 μm CMOStechnology. The circuit is a general purpose test amplifier suitable forthe readout of silicon strip detectors and has already been used in LHCexperiment. For a detailed discussion, reference is made to J. Kaplan,W. Dabrowski, “Fast CMOS Binary Front-End for Silicon Strip Detectors atLHC Experiments”, IEEE Trans. Nucl. Sci., Vol. 52, No. 6, pp. 2713-2720,December 2005. The basic chip consists of 16 channels of chargesensitive amplifiers and analogue output drivers. Each channel comprisesa fast transimpedance preamplifier working with an active feedback loop,wherein one stage of the amplifier-integrator circuit provides 22 nspeaking time, and an output buffer capable to drive up to 10 pF outputload capacitance.

As can be seen from FIG. 5, the preamplifier is based on a classicalcascode configuration with the NMOS input transistor of size 1000 μm/0.5μm optimized for input capacitances in the range of 10 pF. The activefeedback circuit employed in the design offers much lower parasiticcapacitance of the feedback loop and consequently a higher bandwidth ofthat stage, when compared to a resistive feedback loop usinglow-resistivity poly-silicon resistors. The contribution of the activefeedback circuit to ENC noise for a nominal feedback current of 800 nAis about 400 e-, which is still acceptable in view of the expectedseries noise contribution from the input transistor which is loaded withan external capacitance of 10 pF, which is the typical capacitance of anAPD array element. The nominal bias current of the input transistor canbe adjusted between 300 μA to 600 μA, which provides a way to optimizenoise against the power consumed by this stage, as will be explainedbelow.

The following integrator stage comprises a voltage amplifier consistingof two cascaded common source amplifiers enclosed with resistivefeedback. The whole preamplifier-shaper circuit has a gain of about 28mV/fC and a peaking time of 22 ns. The dynamic range of the whole chainis in the range of 800 mV which is equivalent to 30 fC signal charge.Note that in principle the dynamic range of the amplifier in terms ofamplitude response linearity is limited to about 30 fC. However, thepulse width modulation discriminator 54 (described in more detail below)senses input charges that are proportional to a signal pulse shape area.Accordingly, a dynamic range up to 150 fC can be achieved by this chargesensing method.

FIG. 5A shows calculated ENC noise as a function of the inputcapacitance for two different values of the input transistor bias andfor a nominal value of the feedback bias current of 800 nA. As can beseen from FIG. 5A, for a higher input transistor bias I_(drain), asmaller ENC can be achieved. That is, generally, one may have to find acompromise between ENC noise and power consumption. Since the powerconsumption is not of primary importance with regard to the medicalimaging apparatus 28, a bias current of 600 μA is acceptable, andaccordingly, achievable noise levels for typical detector capacitancesof about 10 pF are below 800 e-ENC.

Moreover, since a typical value of threshold applied to the pulse widthmodulation discriminator 54 is in the range of 1 fC, the noise hit ratecaused by the electronics is negligible. Note that in the apparatus 28of the current embodiment, the minimum signal to be detected originatesfrom 100 KeV X-ray photons, which in a typical setup correspond to acharge of about 6 fC, while the PET signal in the same setup correspondsto about 30 fC.

FIG. 6 shows a conceptual diagram of the pulse modulation discriminator54. The discriminator circuit 54 comprises an input stage 62 which isshown in greater detail in FIG. 7. The input stage 62 comprises inputs64 for a current signal received from the preamplifier 52 (cf. FIG. 4A)and inputs 66 for receiving a threshold DC-current signal from athreshold control unit 68. With continued reference to FIG. 6, thediscriminator circuit 54 comprises four differential amplifiers A1 to A4connected in series by which the output of the input stage 62 isamplified. The discriminator circuit 54 further comprises a hysteresisor positive feed-back element 70 for increasing the stability of theoutput signal. Finally, the discriminator circuit 54 comprises an outputstage 72 for shaping the output signal such as to acquire sharp risingand falling edges. The complete discriminator circuit 54 is integratedin monolithic CMOS technology.

In FIG. 7, the input stage 62 is shown in greater detail. As can be seenin FIG. 7, the input stage 62 comprises two symmetrically arrangedcurrent sources 74, 76 connected to transistors M1, M2, respectively. Inaddition, the input stage 62 comprises cascode transistors M3 and M4connected with transistors M1, M2, respectively.

As is readily appreciated by the person skilled in the art, the inputstage 62 of FIG. 6 provides a comparison of the signal current injectedat inputs 64 and the threshold current injected at inputs 66.

It is noted that the discriminator circuit 54 as shown in FIGS. 6 and 7as such is known from the article “NINO: An Ultrafast Low-PowerFront-End Amplifier Discriminator for Time-of-Flight Detector in theALICE Experiment” by F. Anghinolfi, P. Jarron, F. Krummenacher, E.Usenko and M. C. S. Williams, IEEE Transactions on Nuclear Science, vol.51, No. 5, October 2004, pages 1974 to 1978, and from “NINO: Anultra-fast and low-power front-end amplifier/discriminator ASIC designedfor the multigap resistive plate chamber”, F. Anghinolfi, P. Jarron, A.N. Martemyanov, E. Usenko, H. Wenninger, M. C. S. Williams and A.Zichichi, Nucl. Instrum. Methods Phys. Res., A 533 (2004) 183-187.However, this discriminator has been employed in a very differentcontext of time-of-flight detection of charged particles having energiesin the GeV range in high energy physics experiments but not for thedetection of photons and, in particular, not in the context of medicalimaging. In particular, when the discriminator circuit 54 was used inhigh energy physics experiments, the energy of the detected particleswas not encoded and could not even have been encoded, since in the priorart experimental setup the particles to be detected only transmitted apart of their energy to the detector and the information processed forthe experiment comprised only binary if-data (hit or no-hit).

The upper diagram of FIG. 8 shows two simplified waveforms of detectionsignals 80, 82 inputted into the pulse modulation discriminator 54. Theupper signal 80 corresponds to the detection of a photon of higherenergy and the lower signal 82 corresponds to the detection of a photonof lower energy. The signal shape is a triangular shape approximatingthe typical response of a scintillator. That is, typically thescintillator has a characteristic rise time which is largely independentof the energy of the photon and a consecutive exponential decrease at atime constant on the order of 40 ns.

Also in the upper diagram, the discriminator threshold described withreference to FIGS. 6 and 7 above is shown. From the signals 80 and 82and the discriminator threshold, the pulse modulation discriminator 54generates respective pulses shown in the lower diagram as 84 and 86,respectively. The leading edge of the pulse signals 84, 86 correspondsto the time the signal 80, 82, respectively first exceeds thediscriminator threshold, and the trailing edge of the pulse signal 84,86 corresponds to the time the detection signal 80, 82, respectivelydrops below the discriminator threshold. As is clear from FIG. 8, thehigher the amplitude of the signal, or in other words, the higher theenergy of the detected photon, the wider the width of the pulse signal84, 86 will be. Accordingly, the width of the pulse signals 84, 86 canbe used for encoding the energy of the corresponding photon.

As can be seen from FIG. 8, the leading edges of the pulse signals 84,and 86 are not exactly coinciding with one another and in particular,are not coinciding with the true beginning of the signals 80 and 82.Instead, there is a delay corresponding to the time needed for thesignals 80, 82 to reach the discriminator threshold. This delay isusually called “time-walk”. Consequently, while the leading edge of thepulse signals 84, 86 do encode the time-stamp, said time-stamp should becorrected by the time-walk to provide high precision time measurements.As can be discerned from FIG. 8, the time-walk can be approximated bythe following simplifying equation:

${{time}\text{-}{walk}} = {{rise}\mspace{14mu}{time}\mspace{14mu}\frac{{discriminator}\mspace{14mu}{threshold}}{{signal}\mspace{14mu}{amplitude}}}$

Since the pulse width is related to the signal amplitude (in fact, for aconsiderable region it is almost proportional thereto), the time-walkcan be determined from the pulse width. Accordingly, from the pulsesignals 84 and 86 alone, a very precise time-stamp and a precisemeasurement of the energy can be obtained. Moreover, it is noted thatsuch a pulse signal can be easily and rapidly processed, for example bya TDC as described below, allowing for a real time processing of thedata and “on-line” generation of medical images.

With continued reference to FIG. 8, there will be a jitter related tothe time uncertainty with which the signal 80, 82 crosses the thresholddue to noise superimposed on the signal. Timing jitter is given by thegeneral formula σ_(jitter)=σ_(n)/|dV/dT/| where σ_(n) is the noisesuperimposed on the pulse and dV/dT is the slope of the pulse at thepoint where it crosses the threshold. In the situation of FIG. 8, thejitter equation therefore reads as follows: σ_(jitter)=σ_(n)

$\frac{{rise}\mspace{14mu}{time}}{{signal}\mspace{14mu}{amplitude}} = \frac{{rise}\mspace{14mu}{time}}{{signal}\text{-}{to}\text{-}{noise}\mspace{14mu}{ratio}}$Consequently, with a rise time of 1 ns and a threshold to amplituderatio of 5 the time-walk is 200 ps and the jitter would be 20 ps with asignal-to-noise ratio of 50. Such a small jitter allows for very precisemeasurements, well-sufficient for time-of-flight measurements.

FIG. 9 shows three output waveforms 88, 90 and 92 obtained in experimentwith the pulse modulation discriminator 54 of FIG. 6. In thismeasurement, the input of the pulse modulation discriminator 54 has beenstimulated with electronic pulses of 20 fC (curve 88), 80 fC (curve 90)and 240 fC (curve 92), respectively.

As can be seen from FIG. 9, the pulse width is directly related to thesize or charge of the stimulation signal. Also, from FIG. 9 it can beseen that the pulse modulation discriminator even works with much fastersignals than the ordinary scintillation signal as shown in FIG. 8.Clearly, the time resolution obtainable with the pulse modulationdiscriminator is well in the picosecond range allowing for precisetime-of-flight measurements.

FIG. 10 shows measurements of the correction of the time-walk based onthe discriminator pulse width indicating a timing precision easily inthe range of the precision required for a TOF-PET-scanner system.

FIG. 11 is a timing diagram similar to the timing diagram of FIG. 8. Inthe upper diagram of FIG. 11, two detection signals 94, 96 are shownwhich are due to X-ray photons hitting the same detector segment.Comparing the X-ray signals 94 or 96 with the PET-signal 80, it can beseen that the PET-signal 80 is larger, since the PET-γ-photon has anenergy of 511 keV which is considerably higher than the energy oftypical X-ray photons used in CT having about 100 keV. However, thediscriminator threshold of the pulse modulation discriminator 54 isadjusted such as to generate pulses for both, PET and CT events usingthe same discriminator circuit, the pulses only differing in theirrespective lengths.

In the situation shown in FIG. 11, the signals 94 and 96 overlap intime. This overlap or so called “pulse pile-up” frequently occurs atlarger X-ray intensities. As can be seen from the lower diagram of FIG.11, such an overlap of two X-ray signals will lead to a pulse signalwhich is wider than the pulse signal corresponding to a single X-rayphoton. An example of the pulse signal due to two overlapping X-raysignals is also shown in FIG. 14. Overlapping pulse signals will occurif the width of the pulse signal corresponding to a single X-ray photonis larger than the delay time between two consecutive photons hittingthe same detector segments.

In FIG. 12, a pulse pile-up filter 98 is shown. The pulse pile-up filter98 comprises three basic gate-delay circuits 100, 102 and 104 which areconnected in parallel. Each of the basic gate-delay circuits 100, 102,104 comprises an AND-gate to which a fraction of the pulse signal and adelayed fraction of said pulse signal is fed, which delayed fraction isdelayed by a predetermined time-delay. The basic gate-delay circuit 100comprises a delay element providing for a time-delay corresponding to anexpected width of the pulse signal for a single X-ray photon. Gate-delaycircuits 102 and 104 each comprise a delay element providing for delayscorresponding to an expected pulse width of the pulse signals for two orthree overlapping X-ray detection signals, respectively. Accordingly, byusing the pulse pile-up filter 98, two or three X-ray photonsoverlapping in time can be detected and distinguished. The outputs ofthe gate-delay circuits 100, 102, 104 are inputted to a pile-up decoder106 for correcting a counter by one, two or three counts, depending onthe output of the gate-delay circuits 100, 102 and 104.

FIG. 13 shows an example of the PET-CT-filter 56 that has been showngenerally in FIG. 4A. The PET-CT-filter circuit 56 has an input 108 forreceiving a signal from the pulse modulation discriminator 54. Thefilter 56 comprises a CT branch 110 and a PET branch 112. In the CTbranch, a gate-delay circuit is used to check whether the inputted pulsesignal has a width longer than a predetermined delay time CTDmax. CTDmaxmay for example correspond to a pulse with corresponding to a photonhaving an energy of 120 keV. If the output of the AND-gate in the CTbranch 110 is positive, this means that the detected pulse correspondsto a photon having an energy higher than said 120 keV and it cantherefore not be an X-ray photon from the CT-X-ray source 48 of theapparatus 28 shown in FIG. 4. Consequently, this event is not counted bythe corresponding CT counter.

In the PET-branch of the filter 56, it is checked whether the pulsewidth is longer than some predetermined delay PETDmax and longer thansome minimal pulse width Dmin. With this filter, it can be checkedwhether the energy of the corresponding photon is within the rangeexpected for PET-γ-photons. If the energy is within the expected range,the pulse signal will be channeled to the TDC 42 (see FIG. 4).

In FIG. 15, a block diagram of a multichannel TDC circuit 42 is shown.For simplicity, in the multichannel TDC 42 of FIG. 15, only twochannels, 114 and 116 are indicated. Each TDC channel 114, 116 receivespulse signals from the corresponding pulse modulation discriminator viaa respective input. Each TDC channel 114 has means 118, 120 for leadingand trailing edge capture, respectively. The leading edge of the pulsesignal is time-walk corrected and decoded to yield a time-stamp. Thedifference between leading and trailing edge provides the energy orsignal amplitude information. The time-stamp and energy information isstored in a local FIFO register 122. The basic TDC measurements of theleading and trailing edges of the pulse signal are performed in a fullyclock synchronized fashion based on basic clock synchronized counterswith sufficient dynamic range to cover the full period of datacollection. A clock driven fine time interpolator circuit based on phaselocked loops and/or delay locked loops driven from the same clockreference is used to obtain the required time resolution. Thisguarantees a continuously self-calibrating and very stable timemeasurement with very high time resolution (on the order of 10 ps) andvery high operational stability.

As is schematically indicated in FIG. 15, multiple TDC channels can usethe same time counters and fine interpolation units while usingindependent time capture circuits and a local FIFO 122. The timemeasurements from independent channels can be derandomized and mergedinto larger storage memories or FIFOs with an associated channelidentifier and finally be transferred to the subsequent imageprocessing.

FIG. 16 shows a block diagram of the PET readout architecture employingTDCs with local time measurements based on global time reference. InFIG. 16, again an annihilation site 48 is indicated from which γ-photons50 are emitted. Also shown are four exemplary detection elements 34 andcorresponding readout circuitry 40 and TDCs 42.

Each local TDC 42 is connected with a global reference clock 60. Aglobal clock reference is distributed across the system with very smalljitter together with a global clock synchronous reset signal defining aglobal zero time reference plus a time scale given by the repeated clockcycles.

Each TDC uses the received clock reference to measure the arrival timeof the input signal edges. Consequently, only channel time offsetparameters need to be determined from a simple global timing calibrationto perform coincidence determinations and time-of-flight measurementsacross channels throughout the system.

As mentioned above, the present invention allows for a very precisetime-stamp measurement of photons simultaneously with a precisemeasurement of the photon energy. The energy measurement has a number ofimportant advantages. Generally, as has been mentioned above, by aprecise energy measurement, the signal-to-noise ratio can be enhanced.First of all, due to the energy measurement, the envisaged photon eventscan be distinguished from background. Also, a precise energy measurementfurther allows to reconstruct Compton scattering of a γ-photon, that is,it allows to measure the energy of a photon and a corresponding electronafter Compton scattering and to reconstruct the original photontherefrom.

Also, as has been explained with reference to FIG. 13, due to themeasurement of the energy, it is possible to for example discriminatebetween X-ray photons and PET photons using the same front-end detectionarchitecture, allowing for a very simple combined PET-CT apparatus.However, the energy discrimination could for example also be used todistinguish between photons of different wavelengths originating fromdifferent types of radiopharmaceuticals which are applied to a patient'sbody simultaneously. In such an embodiment, it is possible toinvestigate the location and concentration of different types ofradiopharmaceuticals in only a single scan.

In FIGS. 17 and 18, the results of benchmark experiments related to theenergy encoding are shown. In both diagrams, the horizontal axis denotesthe pulse width as generated by the discriminator circuit 54 and thevertical axis represents the number of counts or detected events in agiven period of time.

In particular, FIG. 17 is a histogram plot of the spectrum obtainedusing a ⁵⁷Co radioactive source which has a photo peak located at anenergy of 122 keV and is therefore a good benchmark for CT X-ray. Thediagram of FIG. 18 represents a histogram plot of the spectrum obtainedwith a ²²Na radioactive source with one photo peak located at an energyof 511 keV, which is a good reference for the PET γ-photon, and a secondpeak at 1213 keV. The plots represent experimental raw data without anycorrection and linearization. In order to get a correct energy scale, anoffset of about 20 ns would be subtracted and a signal linearizationshould be applied. FIGS. 17 and 18 demonstrate that in fact a veryprecise energy measurement can be performed with the apparatus of theinvention.

Although a preferred exemplary embodiment is shown and specified indetail in the drawings and the preceding specification, these should beviewed as purely exemplary and not as limiting the invention. It isnoted in this regard that only the preferred exemplary embodiment hasbeen shown and specified, and all variations and modifications should beprotected that presently or in the future lay within the scope ofprotection of the invention as defined in the appending claims.

LIST OF REFERENCE NUMBERS

-   10 APD array-   12 preamplifier-   14 shaper-   16 A/D converter-   18 FIFO buffer-   20 coincidence processor-   22 current preamplifier-   24 gated integrator-   26 A/D converter-   28 combined PET-CT-system-   30 detector ring-   34 detector element-   36 scintillator array-   38 photodetector array-   40 front-end-   42 TDC-   44 coincidence processor-   46 PET-CT-reconstruction processor-   48 γ-photon-   52 preamplifier-   54 pulse modulation discriminator-   56 filter-   58 counter-   60 global reference clock-   62 input stage-   64, 66 inputs-   68 threshold control unit-   70 feed-back element-   72 output stage-   74, 76 current sources-   80 signal waveform-   82 signal waveform-   84 pulse signal-   86 pulse signal-   88-92 measured pulse signals-   94, 96 overlapping X-ray signals-   98 pulse pile-up filter-   100-104 gate-delay circuits-   106 pile-up decoder-   108 input to filter 56-   110 CT-branch-   112 PET-branch-   114, 116 TDC channels-   118 leading edge capture means-   120 trailing edge capture means-   122 FIFO

The invention claimed is:
 1. Apparatus for use in medical imaging, saidapparatus comprising a readout circuit having an input for receiving adetection signal corresponding to a photon hitting a radiation detector,characterized in that the readout circuit is adapted to output, inresponse to receiving said detection signal, a pulse signal having aleading edge encoding a time-stamp of said photon and a width encodingan energy of said photon.
 2. Apparatus according to claim 1, whereinsaid readout circuit is adapted for encoding the time-stamp and energyfor one or more types of photons selected among a group consisting of:γ-photons as generated during electron-positron annihilation, X-rayphotons of a wavelength suitable for CT, visible light photons, andγ-photons as emitted by radiopharmaceuticals suitable for γ-cameraimaging or single photon emission computed tomography imaging.
 3. Theapparatus of claim 1 further comprising means for correcting thetime-stamp based on the width of the pulse signal.
 4. The apparatus ofclaim 3, wherein said correcting means are configured to estimate atime-walk of the pulse signal from the width of the pulse signal and tosubtract the time-walk from the time-stamp.
 5. The apparatus accordingto claim 1, wherein said readout circuit is a discriminator circuitconfigured to compare the detection signal to a threshold value, wherethe leading edge of the pulse signal corresponds to the time thedetection signal first exceeds said threshold and the trailing edge ofthe pulse signal corresponds to the time the detection signal dropsbelow said threshold value.
 6. The apparatus of claim 1, wherein thereadout circuit is a monolithic CMOS-device.
 7. The apparatus accordingto claim 1, wherein the radiation detector comprises a scintillatorelement and a photodetecting element, said photodetecting element beingarranged to receive light emitted from said scintillator element.
 8. Theapparatus according to claim 7, wherein said scintillator elementcomprises a Lu₂SiO₅ crystal or a LuYAP crystal.
 9. The apparatus ofclaim 7, wherein said photodetecting element comprises at least oneavalanche photodiode.
 10. The apparatus according to claim 1, furthercomprising a filter for filtering the pulse signal according to itswidth.
 11. The apparatus according to claim 10, wherein the filtercomprises at least one gate-delay circuit comprising an AND-gate towhich a fraction of said pulse signal and a delayed fraction of saidpulse signal is fed, said delayed fraction being delayed by apredetermined time-delay.
 12. The apparatus according to claim 10,wherein the filter comprises a PET-filter configured to pass pulsesignals having a width corresponding to the energy of γ-photons emittedduring electron-positron annihilation.
 13. The apparatus of claim 12,wherein the PET-filter blocks pulse signals having a width correspondingto an energy of less than 350 keV.
 14. The apparatus according to claim10, wherein the filter comprises a CT-filter configured to allow pulsesignals having a width corresponding to the energy of X-ray photonssuitable for CT to be counted by a counter.
 15. The apparatus accordingto claim 14, wherein the CT-filter blocks pulse signals having a widthcorresponding to an energy of more than 250 keV.
 16. The apparatusaccording to claim 1, further comprising a filter for distinguishingpulse signals corresponding to two or more detection signalscorresponding to individual X-ray photons overlapping in time.
 17. Theapparatus of claim 16, wherein the filter comprises a number of ngate-delay circuits connected in parallel, each gate-delay circuitcomprising an AND-gate to which a fraction of said pulse signal and adelayed fraction of said pulse signal is fed, said delayed fractionbeing delayed by a predetermined time-delay, wherein n is an integergreater or equal to 2, wherein the predetermined time-delays of thegate-delay circuits respectively correspond to an expected width of thepulse signal for one, two, . . . , n overlapping detection signalscorresponding to X-ray photons overlapping in time.
 18. The apparatusaccording to claim 1, further comprising a counter for counting thenumber of X-ray photons detected during a predetermined time period. 19.The apparatus of claim 1, further comprising a time-to-digital converter(TDC) connected to receive the pulse signal and configured to measurethe arrival time of both the leading and trailing edge of the pulsesignal.
 20. The apparatus according to claim 19, wherein the TDC isconfigured to store the leading edge time measurement and the pulsewidth in a local channel FIFO register.
 21. Detection assembly for amedical imaging apparatus, comprising a multitude of apparatusesaccording to claim
 1. 22. The detection assembly according to claim 21,wherein some or all of the multitude of detector elements are connectedto a common reference clock.
 23. A combined PET-CT-apparatus comprisingan X-ray generator suitable for CT and a detection assembly according toclaim
 21. 24. Detection assembly for a medical imaging apparatus, and inparticular for a PET or a PET-CT-apparatus, said detection assemblycomprising a multitude of detector elements, each detector elementcomprising an array of radiation detectors providing a number ofchannels for outputting detection signals and a corresponding number ofapparatuses according to claim 1 providing electronic channels forreading out the detection signals.
 25. The detection assembly of claim24, further comprising a control unit configured or programmed fordetecting coincidence of detection signals.
 26. The detection assemblyaccording to claim 25, wherein said control unit is further configuredor programmed to calculate the location of an electron-positronannihilation using a difference in time-stamp of the coinciding events.27. A method of reading out detection signals from a radiation detectorof a medical imaging apparatus, said method comprising the steps of:receiving an analog detection signal from a radiation detector upon aphoton hitting said radiation detector, generating from said detectionsignal a pulse signal having a leading edge encoding the time-stamp ofsaid photon hitting said radiation detector and a width encoding theenergy of said photon.
 28. The method according to claim 27, whereinsaid photon can be of one or more types selected among a groupconsisting of: γ-photons as generated during electron-positronannihilation, X-ray photons of a wavelength suitable for CT, visiblelight photons, and γ-photons as emitted by radiopharmaceuticals suitablefor γ-camera imaging or single photon emission computed tomographyimaging.
 29. The method of claim 27, further comprising a step ofcorrecting the time-stamp based on the width of the pulse signal. 30.The method of claim 29, wherein the correction step includes estimatinga time-walk based on the width of the pulse signal and subtracting saidtime-walk from said time-stamp.
 31. The method according to claim 27, inwhich the step of generating said pulse signal comprises a comparison ofthe detection signal with a threshold value, wherein the rising edge ofthe pulse signal corresponds to the time said detection signal firstexceeds said threshold, and the trailing edge of said pulse signalcorresponds to the time the detection signal drops below said threshold.32. The method according to claim 27, further comprising a step offiltering each pulse signal according to its width.
 33. The methodaccording to claim 32, wherein the filtering comprises directing pulsesignals having a width corresponding to the energy of a γ-photon emittedduring electron-positron annihilation to a TDC.
 34. The method accordingto claim 32, wherein the filtering comprises allowing pulse signalshaving a width corresponding to the energy of an X-ray photon used in CTto be counted.
 35. The method according to claim 32, wherein thefiltering comprises a determination of whether the width of the pulsesignal corresponds to an expected width occurring when the detectionsignal corresponding to two or more X-ray photons overlap in time. 36.The method according to claim 27, comprising a step of receiving thepulse signal at a TDC and measuring the arrival time of both the leadingand the trailing edge of the pulse signal.
 37. The method according toclaim 36, said method further comprising a step of storing the leadingedge time measurement and the pulse width in a local channel FIFOregister.
 38. The method according to claim 27, further comprising astep of coincidence detection of two detection signals corresponding tothe two γ-photons originating from electron-positron annihilation. 39.The method according to claim 38, further comprising a step ofcalculating the location of the electron-position annihilation using adifference in time-stamp of the coinciding events.